1. Field of the Invention
The present invention relates to temporary therapeutic devices to work in conjunction with a diseased or failing heart to satisfy the hemodynamic needs of a patient. More particularly, the invention relates to a ventricular cuff for assisting a heart to pump blood by intermittently applying pressure to at least a portion of the ventricular surface of the heart at predetermined intervals to aid the ventricles in properly contracting.
2. Discussion of the Related Art
The human heart is a very complicated organ that relies on both mechanical and electrical operation in order to properly perform. As with any complicated mechanism, problems can and do arise, and the heart is no exception. For example, over time the electrical pathways in the heart (which sequentially cause the atria and ventricles to contract) may fail, thereby causing the heart to lose its rhythm, which is known as arrhythmia. In that event, the ventricles will contract at improper times, and as a result the output of blood decreases. In addition, in some failing hearts the muscle of the heart no longer contracts the ventricles to a sufficient extent, also resulting in a dangerous reduction in the amount of blood flow.
Numerous attempts have been made to assist these diseased or failing hearts by applying external pressure directly to the heart. One such example is direct manual compression of the heart by a person's hand during open chest cardiopulmonary resuscitation. Often, however, the patient requires cardiac or circulatory support for extended periods of time, such as hours, days, or even weeks, and it is quite difficult for medical personnel to apply a rhythmic pulsating pressure for such an extended period of time. Further, it is difficult if not impossible to apply by hand a uniform compressing force to a significant portion of the exterior ventricle surface of the chamber of the heart. Moreover, the chest should not be opened for extended periods of time because of the increased risk of infection. As such, manual manipulation of the heart is not a solution to the problem in most cases.
To overcome this problem, mechanical devices have been developed to apply external pressure directly to the heart. Some of these devices utilize an inflatable liner that surrounds the heart. For example, U.S. Pat. No. 5,119,804 to Anstadt discloses a cup that is provided with an elastomeric liner. The heart is held in place within the liner, which is cyclically inflated and deflated to apply external pressure to the heart. While this device provides an improvement in hemodynamics for a diseased or failing heart, the device nevertheless suffers from shortcomings, one being the fact that only a fraction of the external fluid pressure that is applied in the cup inlet to displace the liner, which in turn displaces the heart wall, is transmitted to the heart itself to assist in pumping blood. As the liner is inflated and stretched, a transmural pressure is created in the liner. The transmural pressure in the liner is the difference in pressure that is applied to both sides of the liner. In other words, the transmural pressure is the pressure within the liner that is generated by the elastic wall tensions of the liner. As illustrated in FIG. 9 and described in Augmentation of Pressure In A Vessel Indenting the Surface of the Lung, 1987, by Joshua E. Tsitlik, et al. the transmural pressure (P.sub.tm) for a stretched liner is: EQU P.sub.tm =P.sub.tm -P.sub.out =T.sub.1 /R.sub.1 +T.sub.2 /R.sub.2
where the radii R.sub.1 and R.sub.2 are the maximum and the minimum radii of the membrane curvature, respectively, i.e., the principal radii of curvature. The vectors T.sub.1 and T.sub.2 are the elastic wall tensions (the force per unit length) acting along the edges of the surface element.
In practice, as the liner is inflated, because of its axial length limitation, it stretches and bulges radially inwardly. Thus, the transmural pressure of the liner is directed in the radially outward direction (i.e., away from the heart), such that the pressure applied to the heart is less than the pressure applied to the liner. In addition, due to the bulging of the liner, the heart is deformed into a generally hour-glass shape. In other words, the outer central portions of the ventricles of the heart are deformed inwardly from their normally convex shape into concave shapes (i.e., the heart is indented). Thus, there is not a uniform application of pressure to the outer walls of the ventricles. In addition, a transmural pressure of the indented portion of the heart wall is directed in the radially outward direction and, thus, is subtracted from the fluid pressure that is applied by the liner to the outer surface of the heart. Thus, this transmural pressure is also subtracted from the fluid pressure that is applied within the liner. In other words, the heart wall itself is fighting against the externally applied force. Thus, the externally applied force in devices such as that disclosed in Anstadt does not cooperate with the heart's own natural compressive forces during the systolic phase. It actually fights against the heart's natural motion even when the pressure is applied in synchrony with the natural systolic phase of the heart. As a result, the fluid pressure applied within the liner must overcome the transmural pressure created both in the liner and in the heart wall. Therefore, a relatively high pressure must be applied within the liner (e.g., 150-200 mm Hg) to achieve assistance in circulation support.
Thus, the prior art devices suffer from the further shortcoming that they apply pressure to the heart in a nonuniform manner. Such liners are made from a silicone rubber elastomer, which, when inflated, are inherently distended and assume an inwardly convex shape (as shown in FIG. 9 of the '804 Patent). As a result, those devices cause the heart to indent in its center portion, while allowing the heart ventricles to remain expanded at their upper and lower portions. Therefore, the prior art devices inefficiently assist in pumping blood to and from the heart. As a result, substantial pressure needs to be applied to the inner side of the liner (i.e., P.sub.in) to effectuate displacement of the blood from the ventricles. A considerable portion of the pressure that is applied to the inner side of the liner is wasted because, as described above, transmural pressure is created in the liner and the heart wall.
Another shortcoming inherent in the prior art devices results from the fact that relatively high pressures are applied almost exclusively to the central portion of the ventricles' outer surfaces. This causes the heart to deform into an unnatural shape and may even eventually cause trauma (e.g., bruises) to the heart, especially if one of those devices is operated for an extended period of time.
Several prior art devices apply a vacuum pressure to a relatively small area in the lower portion of the cup to prevent the heart from being ejected from the cup during the systolic compressing phase. The application of a relatively large vacuum to such a small surface of the heart, especially in view of the large external force that needs to be applied to perform the holding function, causes further trauma to the heart.